Focused ablative ultrasound therapies, such as high-intensity focused ultrasound (HIFU) thermal therapy and histotripsy, have demonstrated precise surgical destruction of pathological tissues noninvasively. These therapies rely on focused transducers to deliver high acoustic intensity or pressure to a localized region to ablate the tissue of interest. HIFU requires that a sufficient acoustic intensity be delivered to the focal region for a sufficient time to cause tissue necrosis through heating. For histotripsy, high-pressure, short duration acoustic pulses are applied to cause mechanical breakdown of tissue in the focal volume by inciting cavitation clouds or bubbles through boiling. Histotripsy requires focal pressure levels of 10 to >25 MPa peak negative pressure, and the peak positive pressure can exceed 100 MPa.
In order to generate such pressure levels, careful consideration is required for design of the therapy transducers. Transducers are most commonly constructed from piezoceramic or piezocomposite materials, using spherically curved segments which produce ideal focusing. Typical resonant frequencies vary from 500 kHz to 4 MHz depending on the application. Therapy transducers can range from simple single element construction, to phased arrays of several hundred individual elements to facilitate focal steering. Histotripsy transducers used in previous work have been constructed from piezoceramic or piezocomposite elements which are air-backed and contain a single quarter-wavelength matching layer. These transducers must be large compared to the wavelength to have considerable focal pressure gain. This stipulation can cause significant difficulty when constructing very large, curved transducers which must necessarily maintain high accuracy along the curvature of the surface.
While these transducers have been fairly reliable, they have limited geometric specifications, cost $5,000-$50,000, and require between 1.5-6 months for construction. This turnaround makes it difficult to iterate designs and make small changes to optimize the transducers. It is often desirable to form complex geometries for therapy transducers, for integration and alignment of imaging feedback probes, alignment of multiple therapy elements, generation of complex focal patterns, or maximal utilization of an available acoustic window in the body. Iteration of such transducer designs can be costly and time consuming at the research level due to the associated cost and lead time for producing focused piezoelectric elements and machining required for suitable transducer housings.
This disclosure describes novel designs and methods to construct a focused ultrasound transducer using rapid-prototyping. The transducer comprises multiple flat or high f-number piezoceramic or piezoelectric elements housed in a shell containing acoustic focusing lens and acoustic matching layer, all made using rapid-prototyping. Rapid-prototyping is gaining acceptance in engineering practice as a method to evaluate functional and nonfunctional components in research as an alternative to machining. Rapid-prototyping has advantage over subtractive (e.g., machining) or formative (e.g., injection molding) processes in that it is cost-effective, fast, and can produce nearly limitless complexity. This method can be used to construct single focused transducers as well as array transducers.
Common methods of rapid-prototyping include stereolithography, selective laser sintering, fused deposition modeling (FDM), and 3D printing. Materials used in such machines can be conventional and proprietary polymers, elastomers, plaster, ceramic composites, and metals. The accuracy and resolution of the machines is dictated by their technology. For instance, a stereolithography apparatus (SLA) system can produce resolution in all dimensions between 10-100 μm. This accuracy has been found to be precise enough to align ultrasound elements in the low MHz range.